Antennas of the MRI devices are formed in a known manner using radiating elements with variable shapes made of copper, acting either as transmitting coils in resonance with the circuit that powers it, or as relaxation signal receivers, or both of these functions alternately. They surround the patient or only the part of the body to be analysed.
Placed in the permanent longitudinal magnetic field B0 of an NMR device, they receive an electrical excitation by which they produce the magnetic field B1 orthogonal to B0 at the resonant precession or relaxation frequency (also called the Larmor frequency) of the nuclei of the studied atoms that are located in the magnetic field B0 and/or pick up the radiofrequency (RF) signal at the same resonant frequency of nuclei that were temporarily submitted to the magnetic field B1.
Remember that magnetic spin moments of hydrogen atom nuclei will progressively align under the effect of the static magnetic field B0 along a direction initially parallel to the magnetic field B0 and will result in global magnetisation along the direction of the field B0 called the longitudinal direction z.
When an excitation is applied in pulse form, in other words a radiofrequency field oscillating at the Larmor frequency with a magnetic component denoted B1, perpendicular to B0, the magnetic spin moments will resonate and progressively separate from this longitudinal axis z to move to an inclination angle denoted FA relative to this initial axis describing a movement called a precession movement. Therefore the radiofrequency field B1 can “tilt” the magnetic spin moments at an inclination angle FA relative to the direction of the field B0.
When excitation is interrupted, the magnetic spin moments that separated from their initial axis return to their equilibrium position, in other words the z axis, without stopping rotating. This return to equilibrium is called relaxation. It is then possible to measure this spin rotation movement in the form of a very weak radiofrequency field picked up by the antenna, the picked up radiofrequency field having the same frequency as the exciter radiofrequency field, in other words the Larmor resonant frequency.
Antennas are designed to radiate the pulse magnetic field B1, or to receive signals generated by relaxation of nuclei and preferably to perform these two functions at successive times.
In particular, antennas used for examination of a part of the body and particularly the head, are antennas functioning in near magnetic field, in other words close to radiating physical elements.
Array type antennas have been developed to operate at frequencies of more than 128 MHz. They are formed by a distribution of a plurality of resonators, usually between 8 and 32, acting as transmitters and/or receivers and that are distributed around the sample to be measured.
Each resonator has a specific channel for emission and/or a specific channel for reception of the radiofrequency signal.
In reception, each resonator can thus produce an image of the anatomic region facing which the image is located. The different images are then combined by algorithms to form the final image.
Operation of array type antennas is characterised by non-uniformity of the radiofrequency magnetic fields emitted or received by a single resonator: B1+ in emission and B1− in reception. The quantity of field B1+ corresponds to circular polarisation of the magnetic field rotating in the same direction as the nuclear spins used for imaging. By opposition, the field quantity B1− corresponds to the circular polarisation of the magnetic field that rotates in the inverse direction and that characterises the sensitivity in reception.
In emission, non-uniformity of B1+ results in the appearance of shaded zones or zones with artificial contrast on the image that are difficult to interpret. An array antenna formed by a plurality of resonators are used to overcome this, specifically to:
make B1+ directly uniform;
make the inclination angle uniform.
This compensation for non-uniformity of B1+ is more efficient when the number of resonators in an array antenna is high.
In reception, a larger number of resonators provides a more uniform global reception profile with an increase in the signal-to-noise ratio as described in document [1] (Roemer, P. B. et al. (1990), The NMR phased array, Magnetic Resonance Medicine, 16: 192-225). This increase in the signal-to-noise ratio may be used profitably to increase the resolution of the image or to reduce the acquisition time using an acceleration method that uses the differential sensitivity between the resonators due to their construction or their distribution around the sample.
In brief, the improvement of the capacities and performances of an array antenna depends on an increase in the number of resonators. This approach is more efficient if the size of each resonator remains unchanged. The result is necessarily a reduction in the distance between adjacent resonators that causes an increase in mutual coupling between adjacent resonators. For the same emission efficiency (or reception sensitivity), the increase in mutual coupling is observed by an increase in the transmission coefficient of the radiofrequency signal between the terminals (or radiofrequency ports) of adjacent resonators.
High mutual coupling has three main disadvantages:
an adjustment (i.e. frequency or impedance matching) which is difficult for each resonator because the adjustment of a resonator depends on the adjustment of adjacent resonators;
a loss of efficiency in emission because some of the injected power is dissipated in the circulator load or in the internal resistance of power amplifiers that supply power to adjacent resonators;
an increase in the noise on reception because each resonator receives noise specific to adjacent resonators.
A first analysis based on the circuit theory disclosed in document [2] (Lee, R. F. et al. (2002), Coupling and decoupling theory and its application to the MRI phased array. Magnetic Resonance Medicine, 48: 203-213), can be used to propose a solution for the mutual coupling problem. Efficient decoupling is possible when the characteristics of a resonator can be modelled precisely. However, the difficulty encountered in applying this theory is the result of the strong interaction at high frequency of resonators with the sample to be imaged placed in the antenna, and that modifies model parameters.
Two families of array type antennas are known:
antennas with linear resonators formed from straight copper strips inserted inside an insulating body using the microstrip technique, and usually incorrectly referred to as “antenna strip-line”, and
antennas comprising loop resonators, often formed from a copper strip glued onto a frequently flexible insulating body applied directly onto the sample to be analysed.
Example embodiments have been described in document [3] (Adriany, G, et al. (2005), Transmit and receive transmission line arrays for 7 Tesla parallel imaging, Magnetic Resonance in Medicine 53: 434-445).
Due to the geometric configuration of resonators, mutual coupling is weaker between two adjacent linear resonators than between two adjacent loop resonators, for a given number of resonators in a given space. Consequently, the mutual coupling problem is more severe for antennas comprising loop resonators.
Document [1] describes a method for decoupling loop resonators that consists of superposing loop resonators. Such a solution can lead to a complex configuration that is difficult to implement in practice with patients.
A second known decoupling method consists of adding a passive circuit between two adjacent resonators as described in document [4], (Birl, M. et al. (2005), Surface Coils Decoupling Means for MRI Systems, U.S. Pat. No. 6,927,575 B2).
A third method overcomes mutual coupling problems by combining two resonator technologies, namely linear resonator and loop resonator and using linear resonators in transmission and loop resonators in reception, the resonators operating simultaneously. However, such a configuration cannot optimise the number of resonators in transmission and in reception, therefore there is a loss of efficiency of the antenna. When loop resonators are only used in reception, decoupling can also be achieved by a judicious choice or optimum transformation of the impedance of preamplifiers, as described in document [5] (Yoshida Masaru et al. (2001), RF Coil Array with Reduced Intercoil Coupling, European Patent EP1122550).
If linear resonators are used only in reception, decoupling can also be done using capacitors as described in document [6] (Vaughan, J. T. et al. (2011), Coil Element Decoupling for MRI, US Patent US2011/0312499).
However, none of the described solutions is satisfactory for a Magnetic Resonance Imaging (MRI) application with a very high magnetic field. Furthermore, the disclosed solutions are complex to implement and expensive and require the addition of additional components and electrical circuits.